The present invention realtes to a pulse Doppler measuring apparatus and more in detail to an apparatus for measuring the speed of a body, in particular the blood flow speed in a living body, by using ultrasonic pulses.
When ultrasonic wave is sent to a living body, the frequency of reflected wave is shifted by the Doppler effect due to movement of red blood cells. For example, in FIG. 1A, when ultrasonic pulses are sent to blood cells 4 in a heart 2 with a constant interval T by driving a ultrasonic transmitter/receiver 20 by means of a wave transmitting section 10, in the case where the blood cells 4 move towards the ultrasonic transmitter/receiver 20, a phase difference is produced in the reflected wave coming from the blood cells, which is received by the ultrasonic transmitter/receiver 20 and detected by a receiving section 30. At this time, denoting the blood flow velocity by v, the displacement of the blood cells in this period of time T is given by vT. The phase difference .DELTA..theta. between the preceding reflected wave and the reflected wave at this time is given by: ##EQU1## where k=2.pi./.lambda. .lambda.: wavelength of transmitted ultrasonic wave
f.sub.d : Doppler frequency of moving object PA1 f.sub.d : Doppler frequency of moving object PA1 where EQU A=+2.pi. for .DELTA..theta..sub.n+1 -.DELTA..theta..sub.n &lt;-.pi.
as indicated in FIG. 1B. Using the relation expressed by Eq. (1), it can be known that the phase difference .DELTA..theta. is proportional to vT.
Consequently it is possible to obtain a phase vector of the reflected wave by mixing the reflected wave with a reference wave in a phase measuring section 40, to calculate .DELTA..theta. and obtain further the blood flow velocity v in a velocity calculating section 60 from the phase vector, which has passed through an MTI filter 50. However, as indicated in FIG. 1C, there is a restriction given by, EQU .vertline..DELTA..theta..vertline.&lt;.pi. (2)
on the phase difference .DELTA..theta., which can be measured (because the principal value of the inverse tangent trigonometrical function is taken). Refer to 1978 Ultrasonic Symposium Proceedings IEEE Cat. pp. 348-352.
In an ultrasonic diagnostic apparatus used at present a pulse Doppler blood flow meter is incorporated. This is an apparatus, by which pulsed ultrasonic wave is sent to a body and the blood flow velocity is obtained by using the frequency shift produced by the Doppler effect. The pulse Doppler method has a merit that it is possible to measure the blood flow in a specified part in a body with respect to the continuous wave Doppler method. However, on the other hand, it has a problem that it is not possible to measure any high velocity blood flow exceeding the Nyquist frequency, restricted by the pulse repetition frequency.
Taking diagnosis of heart disease as an example, the measurable limit velocity therefor is about 1 m/sec. On the other hand, in the case where there is abnormal blood flow, the highest velocity thereof is about 5 m/sec. Consequently it is strongly desired by medical doctors to develop a blood flow meter, in which no aliasing takes place.
In the case where the blood flow velocity is large, values different from true values for the velocity and the direction of the blood flow are outputted. This takes place, when a true phase difference .DELTA.'.theta. corresponding to the true velocity exceeds .+-..pi.. The relation between the true phase difference .DELTA.'.theta. and detected phase difference .DELTA..theta. is given by; ##EQU2##
Denoting the pulse repetition frequency by f.sub.s, using Eq (1), the limit of the measurable blood flow is given by: ##EQU3## where f.sub.s =1/T T: interval of transmitted pulses
Accordingly, the limit of the blood flow velocity measurable by the pulse Doppler method is determined by the pulse repetition frequency f.sub.s. Here, according to the condition expressed by Eq. (3), if the pulse repetition frequency f.sub.s is increased, i.e. if the pulse interval T is decreased, it is possible to enlarge the measurable domain of the blood flow velocity.
However, if the pulse interval T is decreased, since the measurable depth is decreased, depending thereon, it is difficult to measure the high velocity blood flow in a deep part. Between the measurable highest blood flow velocity F.sub.m and the depth L.sub.m the following relationship is valid; EQU L.sub.M .multidot.F.sub.M .ltoreq.C/4
where C represents the sound velocity in body.
That is, as far as the aliasing as described previously is produced, it was an important problem how to overcome this problem that there is the relation between the measurable highest blood flow velocity and the depth.
Therefore, in order to produce no aliasing, a method, by which variations are given to the pulse interval (refer to U.S. Pat. No. 4,751,847) and a so-called acceleration method (refer to U.S. Pat. No. 4,884,448) have been proposed. The present invention is an invention, which develops further the latter.
The acceleration method is an algorithm contrived in order to measure precisely a high velocity blood flow (without aliasing), which algorithm is illustrated in FIGS. 2A, 2B and 2C. Hereinbelow the principle thereof will be explained.
According to this method, a vector U.sub.n representing a difference between phase differences is obtained at first by effecting a correlation processing of a phase difference vector Y.sub.n and the complex conjugate vector Y.sub.n-1 * of a directly preceding phase difference vector Y.sub.n-1 obtained by the prior art method. Next, .DELTA..DELTA..theta..sub.n is obtained starting from that vector U.sub.n. That is, ##EQU4## Therefore, using U.sub.Rn and U.sub.In in Eq. (4), EQU .DELTA..DELTA..theta..sub.n =tan .sup.-1 (U.sub.In /U.sub.Rn)(5)
is obtained.
Further, apart from the method described above, there are various sorts of method for calculating .DELTA..DELTA..theta.. For example, EQU .DELTA..DELTA..theta..sub.n =.DELTA..theta..sub.n+1 -.DELTA..theta..sub.n +A
and EQU A=-2.pi. for .DELTA..theta..sub.n+1 -.DELTA..theta..sub.n &gt;.pi.
Here, since the principal value of the inverse tangent trigonometrical function is taken for the difference between phase differences, EQU .vertline..DELTA..DELTA..theta..sub.n .vertline.&lt;.pi.
is valid.
According to the principle of the acceleration method, as indicated in FIG. 2A, the directly succeeding phase difference .DELTA..theta.'.sub.n+1 is obtained by adding the principal value .DELTA..DELTA..theta..sub.n of the difference between two phase differences to the current phase difference .DELTA..theta..sub.n. That is, EQU .DELTA..theta.'.sub.n+1 =.DELTA..DELTA..theta..sub.n +.DELTA..theta..sub.n( 6)
is valid. For this reason, even in the case where the limit given by Eq. (3) is exceeded and the aliasing phenomenon takes place, no aliasing appears, as indicated in FIG. 2C, and thus the blood flow velocity can be measured over the present limitation. Further, since the transmitted pulse interval T remains as it is, the depth of the measurement is also held as it is.
However, in the blood flow measurement by this method, in the case where a true value .DELTA..DELTA.'.theta. of the difference between two phase differences, which is corresponds to a true value of acceleration of blood flow, increases so that .vertline..DELTA..DELTA.'.theta..vertline.&gt;.pi., this gives rise to another problem that an erroneous measurement takes place. The relation between the true value .DELTA..DELTA.'.theta. and a detected value .DELTA..DELTA..theta. is given by; ##EQU5## and thus an error of .+-.2.pi. is produced. Here, this phenomenon being called a second aliasing, the phase difference outputted after the occurrence of the second aliasing continues to have e.g. an error of 2.pi. in Eq. (6) and it is difficult to correct it to return to a correct value.
According to a study of the inventors of the present invention, although (A) production of noise due to characteristics of an MTI filter, (B) noise at the measurement (unexpected movement of the transmitter/ receiver, etc.), (c) rapid variations in the blood flow and (d) low S/N ratio were cited as causes of the production of the second aliasing described above, it has been clearly recognized that (A) the characteristics of the filter are the most important cause of the production.
Here the MTI (Moving Target Indicator) filter is a kind of high-pass filters, which cuts the low frequency component included in the signal reflected by a living body, in the case where the blood flow in a heart or a blood vessel is measured. The low frequency component stated above is produced by the wall motion. Specifically, in the case of the heart, the signal level due to movements of the cardinal wall is extremely high with respect to the blood flow signal level and the difference from the signal level on blood flow information to be obtained is as great as 20 to 60 dB (refer to FIGS. 7 and 9). That is, the MTI filter is used for removing this unnecessary information component.
FIG. 3 shows characteristics of an MTI filter generally used for realizing the acceleration method. In the neighborhood of DC of this MTI filter the signal component due to the wall motion is eliminated. However, since the amplitude of the blood flow signal component is reduced at the neighborhood of zero points of ##EQU6## it is can't be helped that the S/N ratio is lowered there. This is true also at ##EQU7## It is because when ##EQU8## is converted into phase, it is 5.pi. and a case where the phase difference .DELTA..theta. is obtained upto a value 5 times as great as usual one is presumed that data are indicated in the figure up to ##EQU9##
When the S/N ratio is lowered in the neighborhood of such zero points, the signal is apt to include errors, and even in the case where there are no important variations in the blood flow velocity, it can happen that .vertline..DELTA..DELTA.'.theta..vertline. is greater than .pi. and that the secondary aliasing is produced.